Biocompatibility and degradation studies of poly(L-lactide ...
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Biocompatibility and degradation studies ofpoly(L-lactide-co-trimethylene carbonate) copolymers as
cardiac occludersLaishun Xi, Yuandou Wang, Feng Su, Qingzhen Zhu, S.M. Li
To cite this version:Laishun Xi, Yuandou Wang, Feng Su, Qingzhen Zhu, S.M. Li. Biocompatibility and degradation stud-ies of poly(L-lactide-co-trimethylene carbonate) copolymers as cardiac occluders. Materialia, Elsevier,2019, 7, pp.100414. �10.1016/j.mtla.2019.100414�. �hal-03093156�
Biocompatibility and degradation studies of poly(L-lactide-co-trimethylene
carbonate) copolymers as cardiac occluders
Laishun Xi,1 Yuandou Wang,2 Feng Su,1,2* Qingzhen Zhu,1 Suming Li3*
1 State Key Laboratory Base of Eco-chemical Engineering, College of Chemical Engineering,
Qingdao University of Science and Technology, Qingdao 266042, China
2 Institute of High Performance Polymers, Qingdao University of Science and Technology,
Qingdao 266042, China
3 Institut Europeen des Membranes, UMR CNRS 5635, Universite de Montpellier, 34095
Montpellier, France
Correspondence to: F. Su (E-mail: [email protected]) and S. Li (E-mail:
Abstract
Poly(L-lactide-co-trimethylene carbonate) (PLT) copolymers were synthesized by ring
opening polymerization of L-lactide (LLA) and trimethylene carbonate (TMC). The resulting
copolymers were characterized by using 1H NMR, GPC, DSC and tensile tests. The
copolymer properties are dependent on the TMC content. In vitro degradation of copolymers
was carried out in pH 7.4 phosphate buffered saline at 37 °C. The results show that
copolymers with higher TMC content are more resistant to degradation. The cytotoxicity and
hemocompatibility of copolymers were evaluated from MTT assay, hemolysis test, dynamic
clotting time and plasma recalcification time. Results indicate that the copolymers present
very low cytotoxicity and good hemocompatibility.
Cardiac occluders were designed and fabricated using 3D printing. In vivo degradation of
occluders was realized by intramuscular implantation in the back of rabbits. The occluders
were almost totally degraded in 120 days. Visual observation and H&E staining analysis
confirmed the good tissue compatibility of occluders. All these findings suggest that
PLLA-TMC copolymers could be promising for potential applications as degradable occluder
material.
Keywords: copolymer; 3D printing; biocompatibility; cardiac occluder; in vitro degradation;
in vivo implantation
1. Introduction
In the past decades, great progress has been made in the treatment of congenital heart
diseases by the use of cardiac occluders. [1-6] Amplatzer occluders are the most widely used
occluders in clinical applications. They are double-disc devices made from Nitinol wires
tightly woven into two flat discs with a connecting waist. Soft polymeric meshes are placed
inside the discs to block the blood stream. These occluders present good short-term curative
effect, but the permanent presence of metallic devices in the heart often leads to various
adverse reactions such as hemolysis, thrombosis, perforation, allergic skin and dermis, and
even neurological complications. [7] Therefore, occluders based on degradable materials have
received more and more attention.
BioSTAR occluders use an acellular porcine intestinal collagen layer to replace non
degradable meshes. Heparin is coated on the device to reduce protein or blood cell deposition
and thrombus formation. But the device cannot be completely resorbed as the framework is
made of traditional metals or alloys. [8] Totally degradable occluders have been developed in
the past years. Duong et al. developed PFO occluders composed of two self-expanding
umbrellas linked by a stem, and 4 spokes to hold the mesh. The device consists of degradable
poly(ɛ-caprolactone) (PCL) and poly(lactide-co-ɛ-caprolactone) (PLC). [9] Zhu et al. [10]
reported Improved Amplatzer occluders made of braided filaments of poly(para-dioxanone)
(PDO) and poly(L-lactide) (PLLA) nonwoven fabric as a barrier film. Recently, implantation
of totally degradable occluders made of PDO and PLLA was successfully performed in a
clinic operation.
PLLA is a degradable polyester widely used in various biomedical applications such as
drug carriers, surgical implants and sutures because of its biocompatibility and mechanical
strength. [11] PLLA exhibits good tensile strength up to 60 MPa. [12] But PLLA is brittle and
highly crystalline, and acidic degradation products can lead to inflammatory reaction. [13,14]
Furthermore, non-integrative devices with different degradation times could lead to defect
recanalization or abscission of components. Hence, it is of major importance to develop
degradable occluders with sufficient strength, good elasticity and integrative design.
The toughness of PLLA based copolymers can be improved by copolymerization of
L-lactide (L-LA) with other monomers, such as 1,3-trimethylene carbonate (TMC), glycolide
(GA), and ε-caprolactone (ε-CL). [15,16] Poly(1,3-trimethylene carbonate) (PTMC) is an
elastomer with a Tg of -15oC. [17] PTMC degrades extremely slowly by pure hydrolysis,
yielding neutral degradation products, i.e., diols and carbon dioxide. In contrast, PTMC
rapidly degrades in vivo by enzyme catalyzed surface erosion. [18] PTMC is largely used as a
softening component of copolymers for applications as suture material, soft tissue engineering
scaffold, and drug carrier due to its flexibility and biocompatibility. [19-21]
PLLA-TMC (PLT) copolymers have been investigated for applications as heart
constructs and nerve regeneration guides, [22,23] cartilage implants, [24] and sustained drug
release carrier [25, 26] In our previous work, copolymers of L-lactide or DL-lactide and TMC
with various compositions and chain microstructures were synthesized. The thermal
properties, degradation behaviors as well as mechanical properties of the copolymers were
studied to evaluate their potential as cardiovascular stent material. [27-29] The results showed
that PLT copolymers with high LLA contents exhibit high tensile strength and high
crystallinity, but slow degradation rate.
In this paper, PLT copolymers with LLA/TMC molar ratios from 75/25 to 60/40 were
synthesized and characterized. The in vitro degradation and biocompatibility of the
copolymers were investigated to evaluate their potential as degradable occluder material.
Occluders of the copolymers were fabricated by 3D printing. The histocompatibility and in
vivo degradation of the printed occluders were investigated. The results are reported herein in
comparison with literature.
2 Materials and Methods
2.1 Materials
1,3-propanediol and L-lactic acid were purchased from Tianjin Kemiou Chemical
Reagent Co., Ltd (China). Dibutyltin dilaurate, diethyl carbonate, antimony trioxide, zinc
powder, sodium metal and stannous octoate (Sn(Oct)2) were obtained from Sinopharm
Chemical Reagent Co., Ltd (China). The used organic solvents were all of analytic grade.
2.2 Synthesis
LLA and TMC monomers were synthesized as previously reported in literature. [30] The
crude products were purified by recrystallization in ethyl acetate or mixture of diethyl ether
and acetone (V: V = 4: 1). PLT copolymers were synthesized by ring-opening polymerization
of LLA and TMC using stannous octoate as catalyst. [31] Taking PLT 75/25 as example, LLA
(32.36 g), TMC (7.64 g) and the catalyst (0.061 g) were added into a round-bottomed flask,
and degassed. The overall comonomers/catalyst ratio was 2000/1. The polymerization was
carried out under vacuum at 130oC for 72 h. The product was dissolved in dichloromethane,
and precipitated in ethanol. Finally the product was thoroughly dried under vacuum at 50oC
for 72 h.
2.3 Film and occluder fabrication
Copolymer films were prepared by solution casting method. [30] The copolymers were
dissolved in dichloromethane at 10.0 w/v %, and the resulting solutions were poured onto a
glass plate. The solvent was evaporated overnight, followed by vacuum drying for 72 h. The
films were then cut into rectangular samples of 10×75×0.2 mm or square samples of
10×10×0.2 mm for tensile tests and in vitro degradation studies, respectively.
A 3D occluder model with dumbbell shape and hollow structure was designed using
CAD software with parameters of the clinically used devices as reference, and processed
using a 3D printer (UN-biomedical 3d printer, Anyprint, China). The printer parameters were
set as follows: print temperature 140°C, nozzle diameter 0.3 mm, discharge speed 0.012 g/s,
printing rate 4 mm/s, and printing layer thickness 0.2 mm.
2.4 Characterization
Proton nuclear magnetic resonance (1H NMR).1H NMR was performed on Bruker
AVANCE III 500 spectrometer operating at 500 MHz, using deuterated chloroform (CDCl3)
as a solvent. Chemical shifts (δ) were given in ppm using tetramethylsilane (TMS) as an
internal reference.
Gel permeation chromatography (GPC). GPC was conducted on a Shimadzu
apparatus (Waters 410) equipped with a refractometer, using tetrahydrofuran (THF) as mobile
phase at a flow rate of 1.0 mL/min. 60 μL of 1.0 mg/mL sample solution were injected for
measurement, and polystyrene standards were used for calibration.
Differential scanning calorimetry (DSC). DSC was performed using a DSC10
instrument (TA Instruments). 5.0 mg of samples were used for each analysis. A first heating
scan was realized from 0oC to 200oC at 10oC/min, followed by fast cooling down to 0oC at
50oC/min, and a second heating scan to 200oC at 10oC/min.
Tensile testing. Tensile tests were carried out on a GT-TCS-2000 Universal tensile
machine at 25oC, using rectangular samples of 10×75×0.2 mm. The inductor load capacity
was 500 N, and the tensile rate was 50 mm/min. All results are the average of triplicate
measurements.
2.5 In vitro degradation
Square samples with dimensions of 10×10×0.2 mm were weighed and placed in vials
containing 0.1 M pH = 7.4 phosphate buffered saline (PBS) at 37°C. 0.01% (w/v) sodium
azide was added to inhibit the growth of bacteria. At predetermined time intervals, samples
were withdrawn from the vials and washed with distilled water. After wiping, the samples
were vacuum dried up to constant weight before analyses.
2.6 Hemocompatibility
50 mg of films were immersed in 10 mL 0.9% saline, and stirred 72 hours to obtain
extract for hemocompatibility test.
Hemolysis. Hemolysis testing was performed according to the ISO 10993-4-2002
standard, using fresh ACD anticoagulated rabbit whole blood composed of 90% of blood and
10% of 3.8% sodium citrate. Distilled water was taken as positive control, and 0.9% saline as
negative control. [32,33] 10 mL of sample extract in a test tube were thermostated at 37oC for
30 min. Then 0.2 mL of diluted blood with blood/saline volume ratio of 4/5 were added into
the tubes, and maintained at 37oC for 1 h. After incubation, the samples were centrifuged at
3000 r/min for 5 min. The optical density (OD) of the supernatants was measured at an
absorbance wavelength of 545 nm by using UV spectrophotometer (GBC Cintra 10e,
Australia). The hemolytic ratio (HR) was calculated from the OD data of the test sample and
controls using the following formula [34]:
HR (%) = [(ODtest - ODnegative) / (ODpositive - ODnegative)] × 100 (1)
Measurements were made in triplicate (n = 3).
Dynamic clotting time. A method similar to the work described by Zhang et al. [35] was
used for dynamic clotting time measurement. In a test tube were added 10 μL of the sample
extract and 10 μL of 0.2 M CaCl2 solution. After 5 min incubation at 37oC, 80 μL of ACD
blood were introduced in the tube. At different time intervals up to 120 min, 20 mL of
distilled water were carefully added along the tube wall, and the supernatant was collected.
The OD value of the supernatant was determined at 490 nm with a microplate reader (Elx800,
BioTek, USA). Saline was taken as the negative control, and 0.2 M CaCl2 solution as the
positive control. The relative clotting time of samples is obtained from the curves of OD
versus time changes. The experiments were repeated five times for each sample (n = 5).
Plasma recalcification time (PRT). The PRT was determined by using Nie’s method.
[36] Fresh ACD anticoagulated rabbit whole blood was added into an ACD tube, and
centrifuged at 3000 r/min for 10 min to collect platelet poor plasma (PPP). 0.1 mL sample
extract was added to a test tube, followed by addition of 0.1 mL PPP. After 2 min incubation
in a water bath at 37oC, 0.1 mL of 0.025 M CaCl2 solution was added. Every 1 or 2 s the test
tube was tilted so as to observe the state of PPP. The clotting time was taken when the plasma
solution no longer flowed in the inclined tube. Silicified and unsilicified glass tubes were
taken as negative and positive controls, respectively. Triplicate measurements were made for
each copolymer sample (n = 3).
2.7 Cytotoxicity
Both sides of copolymer films (30 mg) were exposed to UV for sterilization at room
temperature for 12 h, and then soaked in 5 mL Dulbecco's Modified Eagle Medium (DMEM).
The samples were placed in water bath oscillator at 37oC for 72 h in a sterile environment,
followed by centrifugation at 1200 rpm for 5 min. Finally the supernatant was collected for
MTT experiment.
MTT test is commonly used to evaluate the cytotoxicity of biomaterials in vitro. [34]
L-929 cells (mouse fibroblasts) in the logarithmic growth phase were collected, and diluted
with DMEM containing 10% fetal bovine serum to a cell concentration of 1 × 104 cells/mL.
100 μL cell suspension was added in a well of 96-well plate, and placed in 5% CO2 incubator
at 37oC. After 24 h incubation, the culture medium was removed and twice washed with PBS.
100 μL sample extract was then added. After incubation for 24, 48 and 72 h, the supernatant
of each well was removed and replaced with 20 μL MTT solution at 5 mg/mL and 80 μL
DMEM. The liquid in the well was removed after 4 h incubation. 150 μL DMSO was then
added, and shaked 10 min. Finally, the OD value was measured at 490 nm by using
microplate reader (Elx800, BioTek, USA). DMEM containing 10% fetal bovine serum was
taken as the negative control, and 6.4% phenol solution as the positive control. Triplicate
measurements were made for all samples. The relative growth rate (RGR) was calculated
according to the following formula [37]:
RGR (%) = (ODtest sample / ODnegative control) × 100 (2)
The cytotoxicity is generally noted in 0-5 levels according to the RGR value. Level 0, 1,
2, 3, 4 and 5 corresponds to RGR ≥ 100%, 100% > RGR ≥ 75%, 75% > RGR ≥ 50%, 50% >
RGR ≥ 25%, 25% > RGR ≥ 1% and RGR > 1%, respectively.
2.8 Histocompatibility and in vivo degradation
PLT copolymer was processed into occluder samples by 3D printing. The thus obtained
samples were soaked in 70% ethanol for 24 h, and then twice washed by using aseptic PBS.
15 New Zealand white rabbits were anesthetized by injection of 1 mL/kg of pentobarbital
sodium into the auricular vein. 4 occluders were implanted into the muscle on the back with a
distance of no less than 5 cm between them. The post-operation rabbits were placed in a
feeding environment at a temperature of 20 ~ 25oC and humidity of 40% ~ 60%. At preset
time intervals (10, 30, 60, 90 and 120 days), three randomly selected rabbits were sacrificed.
The occluders were explanted, and the surrounding tissues removed. The tissues were fixed
with 10% formaldehyde solution for 48 h. After rinsing with water, the tissues were
dehydrated with a series of ethyl alcohol solutions of 70%, 80%, 90%, 95% and 100%. Then
the tissues were made into the paraffin embedding section of a thickness of 5 cm for
hematoxylin and eosin (H&E) staining analysis. The inflammatory reaction and fibrocystic
cavity formation were observed under the microscope.
The recovered occluder samples were successively immersed in type II collagenase
solution (2.5 g/L), and trypase solution (2.0 g/L) for 1 h. The samples were then rinsed twice
with deionized water, and vacuum dried for 1 week up to constant weight. Finally the
degraded samples were characterized by using 1H NMR, GPC and DSC.
3 Result and Discussions
3.1 Characterization of copolymers
A series of PLT copolymers with LLA/TMC molar ratios of 75/25, 70/30, 65/35, and
60/40 were synthesized by ring-opening polymerization, using stannous octoate as initiator.
The chemical composition of the copolymers was determined from 1H-NMR spectra as shown
in Figure 1. The signal at 5.20 ppm belongs to the CH protons of main chain LLA units, and
the smaller signal at 5.05 ppm is assigned to the CH protons of LLA linking to TMC units.
[38] The signals of the CH3 protons of LLA are observed in the 1.5-1.60 ppm (b) range. On
the other hand, the signals of the methylene groups of TMC units are detected at 2.03 ppm (d)
and 4.24 ppm (c), respectively. The LLA/TMC molar ratio of the copolymers was determined
from the integration areas of LLA signal at 5.20 ppm and TMC signal at 4.24 ppm. As shown
in Table 1, the composition of the copolymers is close to the feed ratio, which well
corroborates with high conversion of monomers. The average lengths of lactidyl and
carbonate blocks (lLL and lTMC, respectively) in PLT polymers were determined from 13C NMR
as reported in literature. [27] Both lLL and lTMC increase with increase of the corresponding
component’s content. The lLL increases from 5.29 for PLT 65/35 to 7.55 for PLT 75/25,
whereas lTMC decreases from 2.70 to 2.57 in the meantime.
GPC was employed to measure the molar masses and dispersity of PLT copolymers as
shown in Table 1. The Mn varies from 86100 for PLT 60/40 to 113800 for PLT 75/25, and the
dispersity (Ð=Mw/Mn) varies from 2.0 to 1.8, in agreement with rather narrow molar mass
distribution.
5.5 5.0 4.5 4.0 3.5 3.0 2.5 2.0 1.5
O
O
O O
O Op n
CH3
CH3
a
b
a
b
c
d
c
O
ppm
a c d b
Figure 1. 1H NMR spectrum of PLT 70/30 copolymer in CDCl3.
Table 1. Characteristics of PLLA/TMC (PLT) copolymers.
a Determined by 1H NMR.
b Determined by 13C NMR.
c Determined by GPC.
d Determined by DSC.
e Determined by tensile tests.
f Not determined.
The glass transition temperature (Tg) of the copolymers was obtained from DSC (Table
1). All copolymers exhibit only one glass transition temperature, suggesting that the
copolymer has a relatively random chain structure. Comparison of the Tg data shows that the
addition of TMC component leads to Tg decrease from 42.8°C for PLT 75/25 to 34.1°C for
PLT 60/40. This could be assigned to fact that PTMC has lower Tg and higher chain flexibility
Copolymer LLA/TMC a lLLA b lTMC
b Mn c Ð c Tg (°C) d σ (MPa) e Ε (%) e
PLT 75/25 74.6/25.4 7.55 2.57 113800 1.8 42.8 31.5 87.0
PLT 70/30 70.2/29.8 6.42 2.64 108100 1.8 40.8 20.2 269.1
PLT 65/35 66.2/33.8 5.29 2.70 96500 1.9 37.2 18.3 386.1
PLT 60/40 59.4/40.6 - f - f 86100 2.0 34.1 16.5 512.0
as compared to PLLA. [39,40] All the copolymers are amorphous as no melting peak was
detected. According to the literature, [41] PLLA/TMC copolymers can crystallize when the
content of TMC is below 15 mol%. It has been reported that biomaterials with high
crystallinity could lead to late complications such as inflammatory reactions. [18] Thus the
amorphousness of PLT copolymers should be beneficial for uses as degradable occluders.
The mechanical properties are of crucial importance for occluder materials. Dong et al.
[42] developed PLLA-based terpolymers for applications as biodegradable vascular stent
material. Vascular stents should exhibit higher mechanical strength to prevent recoil due to
vessel pressure as compared to occluders, and need to support the vessels for about 6 months.
In the case of occluders, mechanical support is required for 1-3 months.
The tensile strength (σ) and elongation at break (ε) of PLT copolymers are shown in
Table 1. It appears that the composition strongly affect the mechanical properties of
copolymers. PLLA homopolymer has a tensile strength of 55.6 MPa and an elongation at
break of 9%. [41] With introduction of the TMC comonomer, the tensile strength of PLT
copolymers decreases from 31.5 MPa for PLT 75/25 to 16.5 MPa for PLT 60/40, the
elongation at break increases from 87 % to 512 %. The higher the TMC content, the lower the
tensile strength, and the higher the elongation at break. Therefore, introduction of flexible
TMC component into rigid PLLA chains leads to decrease of the tensile strength and increase
of the elongation at break of copolymers.
3.2 In vitro degradation
PLT 75/25, PLT 70/30 and PLT 65/65 were selected for in vitro degradation studies. PLT
60/40 was not considered as its Tg (34.1°C) is below the body temperature.
Mass loss and water uptake. Mass loss refers to the ratio of soluble species produced
during degradation which are dissolved in the medium, and water uptake reflects the content
of absorbed water in the remaining polymers. Water absorption and mass loss were obtained
according to the following equations [32]:
Mass loss (%) = 100(Wi - Wd)/Wi (3)
Water uptake (%) = 100(Ww - Wd)/Wd (4)
where Wi represents the initial weight, Wd the dry weight after vacuum drying, and Ww the
wet weight of samples after degradation.
Figure 2 presents water uptake and mass loss changes of PLT copolymers during
degradation. During the degradation period up to 90 days, 6.5%, 4.9%, and 4.7% of mass loss
are obtained for PLT 75/25, PLT 70/30, and PLT 65/35, respectively. Larger difference was
observed in water uptake profiles. Water uptake attained nearly 15% for PLT 75/25 after 90
days, whereas PLT 70/30 and PLT 65/35 had a water uptake of 11% and 9%, respectively.
These findings could be attributed to the fact that PTMC is more resistant to hydrolytic
degradation, in agreement with the work reported by Zhang et al. [43] The authors observed
that long TMC sequences can be hardly hydrolyzed although PLT copolymers with TMC
content above 20 mol% are amorphous materials.
0 20 40 60 80 100
0
1
2
3
4
5
6
7
8
9
0 20 40 60 80 100
0
2
4
6
8
10
12
14
16
18
Mass l
oss r
ati
o (
%)
Degradation Time (days)
PLT 75/25
PLT 70/30
PLT 65/35
B
Wate
r u
pta
ke r
ati
o (
%)
Degradation Time (days)
PLT 75/25
PLT 70/30
PLT 65/35
A
Figure 2. Mass loss (A) and water uptake changes (B) of PLT 75/25, PLT 70/30 and PLT
65/35 during degradation.
Molar mass changes during degradation. The variation of Mn during degradation is
shown in Figure 3. A molar mass decrease is observed for all samples from the very beginning.
Actually, water penetrates the PLT samples once immersed in the PBS medium. The chains
are cleaved by hydrolysis of ester and carbonate bonds, resulting in the decrease of molar
mass. The degradation rate of PLT 65/35 is the slowest, with Mn continuously decreasing
from 96500 to 56900 at 90 days, equivalent to 59% of the initial value. PLT 70/30 degrades
slightly faster than PLT 65/35. The Mn decreases from 108100 to 61610 at 90 days, equivalent
to 57% of the initial value. PLT 75/25 shows the fastest decrease of Mn among all the
copolymers with 52% remaining molar mass during the degradation period.
It is worthwhile to point out that polymeric materials lose their mechanical strength
when the Mn is below 25000. [18] The ideal mechanical support time for heart occluders is
2-3 months. Therefore, the three copolymers seem to meet the requirement from the viewpoint
of mechanical support during degradation.
0 20 40 60 80 100
50
60
70
80
90
100
Mn
de
cre
as
e r
ati
o (
%)
Time (days)
PLT 75/25
PLT 70/30
PLT 65/35
Figure 3. Molar mass decrease ratio of PLT 75/25, PLT 70/30 and PLT 65/35 during
degradation.
Compositional changes during degradation. The compositional changes of
copolymers during degradation were determined by 1H NMR as shown in Table 2. The
LLA/TMC ratio of PLT 75/25, PLT 70/30, PLT 65/35 decreases from 3.13 to 2.86, from 2.38
to 2.23, and from 1.94 to 1.82 during 90 days degradation, respectively. In other words, LLA
units are preferentially degraded because they are more degradable than TMC units, as
previously reported in literature. [27,44]
Table 2. LLA/TMC ratio changes of PLT copolymers during in vitro degradation.
Time (days) LLA/TMC ratio
PLT 75/25 PLT 70/30 PLT 65/35
0 3.13 2.38 1.94
40 3.08 2.33 1.85
60 2.94 2.30 1.84
90 2.86 2.23 1.82
3.3 Hemocompatibility evaluation
Hemocompatibility is a key property of biomaterials which are used in contact with
blood. Hemolytic ratio, dynamic clotting time and plasma recalcification time were measured
to evaluate the hemocompatibility of copolymers.
Hemolysis. Hemolysis test is commonly used for the screening of medical materials. It is
generally admitted that a material would not cause hemolysis and can be used for medical
applications if the hemolysis ratio is below 5%, and vice versa. Table 3 presents the hemolysis
ratio of the PLT copolymers. The positive control had an absorbance value of 0.707, whereas
the OD value of the negative control group was 0.0033. These values are within the
recommended range of ISO 10993-4. The hemolysis ratio of all samples is well below 5%,
indicating that the copolymers have little effect on the erythrocytes, and are thus safe for
medical applications.
Table 3. Hemolysis data of the PLT copolymers.
Copolymer OD value Hemolysis ratio (%)
PLT 75/25 0.0063±0.0014 0.426±0.085
PLT 70/30 0.0057±0.0006 0.340±0.020
PLT 65/35 0.0084±0.0006 0.582±0.164
PLT 60/40 0.0067±0.0015 0.475±0.105
Negative control 0.0033±0.0006 -
Positive control 0.7070±0.0360 -
Dynamic clotting time. The clotting time allows to evaluate the activation degree of
intrinsic coagulation factors, and is thus used to assess the influence of biomaterials on the
coagulation process. Absorbance-time curves reflect the coagulation trend of the various
samples as shown in Figure 4. The time at which the absorbance equals to 0.1 is generally
taken as the clotting time. [45] The initial coagulation time of PLT 75/25, PLT 70/30, PLT
65/35, PLT 60/40 is estimated to be 28, 33, 35, 36 min, respectively, which is shorter than that
of the negative control (48 min), but much longer than that of the positive control (6 min).
These findings suggest that the copolymers have little effect on the erythrocytes.
0 20 40 60 80 100 120 140 160 180 200
0.04
0.06
0.08
0.10
0.12
0.14 Negative control
PLT 60/40
PLT 65/35
PLT 70/30
PLT 75/25
Positive control
OD
Va
lue
Time (min)
Figure 4. Absorbance-time curves of copolymers in comparison with the controls.
Plasma recalcification time. Plasma recalcification profile allows to measure the time
of clot formation in recalcified blood, and can serve as indicator of the intrinsic coagulation
system. The plasma recalcification time (PRT) of materials is considered to be significant if it
is above 140% that of the positive control (not silicified glass). [46,47] As shown in Figure 5,
the PRT of PLT 75/25, PLT 70/30, PLT 65/35, PLT 60/40 is 279.7±5.4, 322.0±6.9, 342.3±7.3
and 335.8±9.6 s, respectively. The PRT of copolymers is shorter than that of the negative
control (372.6±10.8 s), but much longer than that of the positive control (154.3±9.1 s).
Moreover, the PRT of all copolymers is largely above 140% that of the positive control, thus
indicating that they do not have noticeable effect on the intrinsic coagulation pathway.
0
50
100
150
200
250
300
350
400
450
500
Tim
e (
s)
Posi
tive
contr
ol
PLT 7
5/25
PLT 7
0/30
PLT 6
5/35
PLT 6
0/40
Neg
ativ
e co
ntrol
*
**
*****
Figure 5. Plasma recalcification time of PLT 75/25, PLT 70/30, PLT 65/35 and PLT 60/40
copolymers. ** indicates p < 0.01, *** p < 0.001, and * p > 0.05.
3.4 Cytotoxicity evaluation
L-929 cells (mouse fibroblasts) are a commonly used standard cell line for cytotoxicity
evaluation of biomaterials which are to be used as medical implants and thus in direct contact
with fibroblasts. The cytotoxicity of PLT copolymers was evaluated by using MTT method.
[48] Figure 6 shows the viability of cells after co-culture with copolymer extracts for 1, 2 and
3 days. The L-929 cells’ viability varies with the incubation time, and the number of L-929
cells reaches a maximum at the second day. The RGR value is above 100% for all copolymer
samples, corresponding to a cytotoxicity level of 0 (Table 4). On the other hand, the survival
ratio of the cells in the positive control is very low, and the cytotoxicity level is 4. These
findings show that there was no release of cytotoxic species in the copolymer extracts.
0.0
0.5
1.0
1.5
2.0 Negative Control
PLT 75/25
PLT 70/30
PLT 65/35
PLT 60/40
Positive Control
OD
Va
lue
1d 2d 3d
Figure 6. Effect of PLT copolymers on L-929 cells growth in comparison with controls.
Table 4. RGR values and cytotoxicity levels of PLT copolymers during 3 days’ incubation
with L-929 cells.
Copolymer 1d 2d 3d
RGR (%) Level RGR (%) Level RGR (%) Level
PLT 75/25 115.0±8.7 0 101.7±7.2 0 128.9±12.4 0
PLT 70/30 118.9±5.8 0 101.9±6.3 0 118.8±8.4 0
PLT 65/35 122.7±10.8 0 107.1±5.0 0 126.9±6.8 0
PLT 60/40 117.7±4.7 0 111.2±6.2 0 121.1±11.2 0
Negative Control 100 0 100 0 100 0
Positive Control 8.3±2.7 4 4.8±1.5 4 9.1±1.8 4
3.5 In vivo degradation
PLT 70/30 copolymer was selected for the fabrication of occluder samples by 3D
printing due to its outstanding overall properties. A hollow structure with dumbbell shape was
designed as shown in Figure 7A. The diameter and the total height are both 12 mm. The
thickness of the wall is 0.2 mm. The 3D printed occluder is shown in Figure 7B. The occluder
samples were implanted into New Zealand white rabbit muscle tissue, and removed at 10, 30,
60, 90, and 120 days, respectively. No death occurred during the operation. No abnormality,
redness or infection was observed in the post-operation period.
Figure 7. Design (A) and 3D printed occluder sample (B) of PLT 70/30.
Figure 8. Shape changes during in vivo degradation of the occluders: (A) 10 days; (B) 30 days;
(C) 60 days; (D) 90 days.
The occluder appeared flattened after 10 days’ implantation as shown in Figure 8A. This
phenomenon could be attributed to the fact that the Tg of the copolymer is only slightly
slightly above the body temperature. Thus the copolymer chains have a certain mobility. And
with the muscle pressure and movement, the hollow occluder structure collapsed. At 30 days,
the occluder became more deformed (Figure 8B). And at 60 and 90 days, the occluder
appeared much smaller (Figure 8C, D). Finally at 120 days, the occluder was almost
completely degraded, leaving dispersed debris around the muscle tissue.
B
C D
A
A
B
Table 5 presents the changes of LLA/TMC ratio, Mn and dispersity of occluder samples
during in vivo degradation. The Mn of occluder sample was initially 55970, which was much
lower than the value of 108100 of the copolymer. In fact, the 3D printing at high temperature
led to a strong Mn decrease, in agreement with thermal degradation of the copolymer.
Therefore, 3D printing conditions should be strictly controlled to minimize thermal
degradation. Especially, the copolymer should be thoroughly dried prior to processing. A
constant decrease of Mn is observed during degradation from initial 55970 to 1290 at 90 days.
The dispersity of molar masses remained almost unchanged. On the other hand, the
LLA/TMC ratio decreased continuously from 2.38 initially to 1.92 at 60 days. There findings
are consistent with in vitro degradation data, and confirmed faster degradation of LLA units
than TMC ones in amorphous zones. Surprisingly, a sharp increase of LLA/TMC ratio up to
4.26 was observed at the last stage of degradation, which could be assigned to crystallization
of PLLA segments as shown in Figure 9.
Table 5. Composition and molar mass changes of PLT 70/30 copolymer during in vivo
degradation.
Time (days) LLA/TMC ratio a Mn b Ð
b
0 2.38 55970 1.8
10 2.22 30000 1.7
30 2.07 15500 1.5
60 1.92 10000 1.8
90 4.26 1290 1.9
a Determined by 1H NMR.
b Determined by GPC.
0 20 40 60 80 100 120 0 20 40 60 80 100 120
B
En
do
therm
al
Temperature (°C)
0d
10d
30d
60d
90d
A
En
do
therm
al
60d
90d
30d
10d
Temperature (oC)
0d
Figure 9. DSC thermograms of PLT 70/30 copolymer after 0, 10, 30, 60, and 90 days in vivo
degradation: (A) first heating scan; (B) second heating scan.
It is of interest to follow the morphological changes of the samples during degradation.
PLT 70/30 is initially an amorphous copolymer which only exhibits a glass transition. The Tg
determined at the second heating scan shows a gradual decrease from initial 40.5° to 32.8° at
60 days. Interestingly, a small endothermic peak is observed at 73.0°C with a melting
enthalpy (∆Hm) of 6.8 J/g at 60 days, indicating crystallization of degradation by-products. A
larger melting peak is detected at 104.1°C with a ∆Hm of 38.6 J/g at 90 days. In fact, PLT
70/30 has a blocky chain structure with a lLL of 6.42 (Table 1). Thus low molar mass
LLA-rich segments produced by degradation are able to crystallize at 37°C due to higher
chain mobility as compared to long chains, as previously reported in literature. [27,49,50]
Once crystallized, the PLLA-rich segments became more resistant to hydrolytic degradation.
As a consequence, the LLA/TMC ratio increased at the late stage of degradation. The Tg also
increased from 32.8°C at 60 days to 42.6°C at 90 days because the Tg of PLLA is much higher
than that of PTMC.
3.6 Histocompatibility evaluation
The histological examination at different stages post-implantation is shown in Figure 10.
After 10 days, no edema or congestion was observed around the tissue. New tissue began to
wrap the sample, and it was not easy to distinguish the boundary between the sample and the
tissues. The new tissue is predominantly composed of fibroblasts and macrophage cells. But
neutrophils, lymphocytes and new capillaries were also observed. At 30 days, the occluder
samples were wrapped by a white fiber tissue which formed a loose and rather thick capsule.
The capsule was composed of three layers: a foam cell layer near the occluder consisting of
degraded copolymer debris phagocytized by macrophages, a middle fibroblast layer, and an
outer fibrocellular layer. The three layers were loosely aligned around the occluder.
Capillaries and some lymphocytes were observed inside the capsule structure. No edema or
hyperemia was observed. At 60 days, the capsule wall became globally thinner and more
compact. Thinning of the fibrocellular and fibroblast layers was detected, together with
thickening of the foam cell layer and the presence of come multinucleated and plasma cells.
At 90 days, the fibrous capsule wall turned thinner and denser, and could be easily detached.
The boundaries between the three layers almost disappeared. Similar observations were made
at 120 days. Foam cells and a few lymphocytes were found in the capsule.
Figure 10. Histological examination at different stages after occluder implantation (original
magnification. X200). (A) 10 days; (B) 30 days; (C) 60 days; (D) 90 days; (E) 120 days.
The degradation of PLT copolymer occluders can be divided in two periods. In the first
period, the molar mass gradually decreases without weight loss. Inflammatory response was
detected in the peripheral tissue after 10 days implantation. In the second period from 30 days,
weight loss occurred until total degradation of the occluders. Small crystalline particles were
obtained during degradation of the material. Meanwhile, macrophages were activated,
resulting in the inflammatory reaction. A large number of macrophages englobed the
crystalline particles, and thus yielding a foam cell layer. With the time passing by, the foam
cells number progressively decreased, and so did the inflammatory reaction. In general, the
endothelialization was completed within 3 months of implantation.
Severe inflammatory reactions are not desirable from the histocompatibility aspect.
Nevertheless, moderate inflammatory reaction could be helpful to complete endothelialization
of the occluder. Cardiac endothelial cells would progressively cover the occluder after
implantation, and endothelialization would be completed in 1-3 months. Occluder needs to
fulfil its function of blocking and supporting before endothelialization, should be degraded
soon after complete endothelialization. If the occluder degrades too quickly, the radial support
would quickly decline. Thus the effective support and blocking effect would be deficient. And
accumulation of large amounts of degradation products in a short time could lead to
occurrence of inflammation and endothelial hyperplasia. On the other hand, if the degradation
time is too long, occluders could cause inflammation, and could lead to formation of
complications such as thrombosis. Further studies are underway in our group to optimize the
copolymer properties, the design of occluder architecture, and the matching of the occluders
and the guidewires for potential clinical applications.
4 Conclusion
In this work, a series of high molar mass poly(L-lactide-co-trimethylene carbonate) (PLT)
copolymers were synthesized and characterized. All copolymers exhibit only one glass
transition, and the addition of TMC component leads to Tg decrease. The tensile strength of
copolymers decreases with increasing the content of TMC component, while the elongation at
break increases. In vitro degradation of PLT copolymers with higher TMC content is slower
than that with lower TMC content because TMC units are more resistant to hydrolytic
cleavage. The various PLT copolymers present good hemocompatibility and low cytotoxicity
as revealed by the hemolysis, dynamic clotting, plasma recalcification and MTT tests.
In vivo degradation was realized by implantation of 3D printed PLT 70/30 occluders in
rabbits. Degradation rate was almost completed after 120 days. In particular, crystallization of
PLLA degradation by-products was observed. Visual observation and H&E staining analysis
confirmed the good tissue compatibility of occluders. Therefore, PLLA-TMC copolymers
could be promising for potential applications as degradable occluder material.
Acknowledgment
The work is supported by the Science and Technology Development Plan of Shandong
Province (2018GGX102016) and the 2018 Shandong Province Graduate Education Joint
Training Base Construction Project.
References
[1] M. Nassif, M. Abdelghani, B. J. Bouma, B. Straver, N. A. Blom, K. T. Koch, J. G. P.
Tijssen, B. J. M. Mulder, R. J. De Winter, Historical developments of atrial septal defect
closure devices: what we learn from the past, Expert Review of Medical Devices 13 (6)
(2016) 555-568.
[2] B. Y. Tang, F. Su, X. K. Sun, Q. Wu, Q. S. Xing, S. M. Li, Recent development of
transcatheter closure of atrial septal defect and patent foramen ovale with occluders,
Journal of Biomedical Materials Research Part B-Applied Biomaterials 106 (1) (2018)
433-443.
[3] Z. Jalal, S. Hascoet, C. Gronier, F. Godart, L. Mauri, C. Dauphin, B. Lefort, M. Lachaud,
D. Piot, M. L. Dinet, Y. Levy, A. Fraisse, C. Ovaert, X. Pillois, J. R. Lusson, J. Petit, A. E.
Baruteau, J. B. Thambo, Long-Term Outcomes After Percutaneous Closure of Ostium
Secundum Atrial Septal Defect in the Young A Nationwide Cohort Study,
Jacc-Cardiovascular Interventions 11 (8) (2018) 795-804.
[4] M. F. Lopez, B. Krastins, D. A. Sarracino, G. Byram, M. S. Vogelsang, A. Prakash, S.
Peterman, S. Ahmad, G. Vadali, W. J. Deng, I. Inglessis, T. Wickham, K. Feeney, G. W.
Dec, I. Palacios, F. S. Buonanno, E. H. Lo, M. M. Ning, Proteomic signatures of serum
albumin-bound proteins from stroke patients with and without endovascular closure of
PFO are significantly different and suggest a novel mechanism for cholesterol efflux,
Clinical Proteomics 12 (2015).
[5] C. M. Happel, K. T. Laser, M. Sigler, D. Kececioglu, E. Sandica, N. A. Haas, Single
Center Experience: Implantation Failures, Early, and Late Complications After
Implantation of a Partially Biodegradable ASD/PFO-Device (BioStar (R)),
Catheterization and Cardiovascular Interventions 85 (6) (2015) 990-997.
[6] Y. Y. Huang, J. F. Kong, S. S. Venkatraman, Biomaterials and design in occlusion devices
for cardiac defects: A review, Acta Biomaterialia 10 (3) (2014) 1088-1101.
[7] S. F. Rimoldi, S. R. Ott, E. Rexhaj, R. Von Arx, S. F. De Marchi, R. Brenner, U. Scherrer,
B. Meier, M. Gugger, Y. Allemann, C. Seiler, Effect of Patent Foramen Ovale Closure on
Obstructive Sleep Apnea, Journal of the American College of Cardiology 65 (20) (2015)
2257-2258.
[8] M. J. Mullen, D. Hildick-Smith, J. V. De Giovanni, C. Duke, W. S. Hillis, W. L. Morrison,
C. Jux, BioSTAR Evaluation STudy (BEST): a prospective, multicenter, phase I clinical
trial to evaluate the feasibility, efficacy, and safety of the BioSTAR bioabsorbable septal
repair implant for the closure of atrial-level shunts, Circulation 114 (18) (2006) 1962-7.
[9] D. Duong-Hong, Y. D. Tang, W. Wu, S. S. Venkatraman, F. Boey, J. Lim, J. Yip, Fully
Biodegradable Septal Defect Occluder-A Double Umbrella Design, Catheterization and
Cardiovascular Interventions 76 (5) (2010) 711-718.
[10] Y. F. Zhu, X. M. Huang, J. Cao, J. Q. Hu, Y. Bai, H. B. Jiang, Z. F. Li, Y. Chen, W. Wang,
Y. W. Qin, X. X. Zhao, Animal Experimental Study of the Fully Biodegradable Atrial
Septal Defect (ASD) Occluder, Journal of Biomedicine and Biotechnology (2012).
[11] M. Abu Ghalia, Y. Dahman, Biodegradable poly(lactic acid)-based scaffolds: synthesis
and biomedical applications, Journal of Polymer Research 24 (5) (2017).
[12] S. Venkatraman, T. L. Poh, T. Vinalia, K. H. Mak, F. Boey, Collapse pressures of
biodegradable stents, Biomaterials 24 (12) (2003) 2105-2111.
[13] Y. Ramot, M. Haim-Zada, A. J. Domb, A. Nyska, Biocompatibility and safety of PLA and
its copolymers, Advanced Drug Delivery Reviews 107 (2016) 153-162.
[14] N. Saba, M. Jawaid, O. Al-Othman, An Overview on Polylactic Acid, its Cellulosic
Composites and Applications, Current Organic Synthesis 14 (2) (2017) 156-170.
[15] A. Smola, P. Dobrzynski, M. Cristea, J. Kasperczyk, M. Sobota, K. Gebarowska, H.
Janeczek, Bioresorbable terpolymers based on L-lactide, glycolide and trimethylene
carbonate with shape memory behaviour, Polymer Chemistry 5 (7) (2014) 2442-2452.
[16] A. K. Matta, R. U. Rao, K. N. S. Suman, V. Rambabu, Preparation and Characterization
of Biodegradable PLA/PCL Polymeric Blends, Procedia Materials Science 6 (2014)
1266-1270.
[17] C. Zhang, D. H. Liu, X. W. Zhang, P. Wang, Z. Zhen, J. X. Li, D. X. Yi, Y. Jin, D. Yang,
Design and in vivo assessment of polyester copolymers based on trimethylene carbonate
and epsilon-caprolactone, Journal of Applied Polymer Science 132 (16) (2015).
[18] A. P. Pego, M. J. A. Van Luyn, L. A. Brouwer, P. B. van Wachem, A. A. Poot, D. W.
Grijpma, J. Feijen, In vivo behavior of poly (1, 3-trimethylene carbonate) and copolymers
of 1, 3-trimethylene carbonate with D, L-lactide or ϵ-caprolactone: Degradation and
tissue response, 67 (2003) 1044-1054.
[19] K. Noorsal, M. D. Mantle, L. F. Gladden, R. E. Cameron, Degradation and drug-release
studies of a poly(glycolide-co-trimethylene carbonate) copolymer (Maxon), Journal of
Applied Polymer Science 95 (3) (2005) 475-486.
[20] B. Y. Yan, Z. P. Zhang, X. P. Wang, Y. F. Ni, Y. S. Liu, T. Liu, W. P. Wang, H. Xing, Y.
Sun, J. Wang, X. F. Li, PLGA-PTMC-Cultured Bone Mesenchymal Stem Cell Scaffold
Enhances Cartilage Regeneration in Tissue-Engineered Tracheal Transplantation,
Artificial Organs 41 (5) (2017) 461-469.
[21] H. Zhao, Y. L. Wang, J. R. Peng, L. Zhang, Y. Qu, B. Y. Chu, M. L. Dong, L. W. Tan, Z. Y.
Qian, Biodegradable Self-Assembled Micelles Based on MPEG-PTMC Copolymers: An
Ideal Drug Delivery System for Vincristine, Journal of Biomedical Nanotechnology 13 (4)
(2017) 427-436.
[22] S. Mukherjee, C. Gualandi, M. L. Focarete, R. Ravichandran, J. R. Venugopal, M.
Raghunath, S. Ramakrishna, Elastomeric electrospun scaffolds of
poly(l-lactide-co-trimethylene carbonate) for myocardial tissue engineering, Journal of
Materials Science-Materials in Medicine 22 (7) (2011) 1689-1699.
[23] C. Marchesi, M. Pluderi, F. Colleoni, M. Belicchi, M. Meregalli, A. Farini, D. Parolini, L.
Draghi, M. E. Fruguglietti, M. Gavina, L. Porretti, A. Cattaneo, M. Battistelli, A. Prelle,
M. Moggio, S. Borsa, L. Bello, D. Spagnoli, S. M. Gaini, M. C. Tanzi, N. Bresolin, N.
Grimoldi, Y. Torrente, Skin-derived stem cells transplanted into resorbable guides provide
functional nerve regeneration after sciatic nerve resection, Glia 55 (4) (2007) 425-38.
[24] N. Andronova, A. C. Albertsson, Resilient bioresorbable copolymers based on
trimethylene carbonate, L-lactide, and 1,5-dioxepan-2-one, Biomacromolecules 7 (5)
(2006) 1489-95.
[25] K. Jelonek, J. Kasperczyk, S. M. Li, P. Dobrzynski, B. Jarzabek, Controlled
poly(L-lactide-co-trimethylene carbonate) delivery system of cyclosporine A and
rapamycine - the effect of copolymer chain microstructure on drug release rate,
International Journal of Pharmaceutics 414 (1-2) (2011) 203-209.
[26] K. Jelonek, B. Kaczmarczyk, J. Jaworska, M. Pastusiak, M. Sobota, P. Dobrzynski, J.
Kasperczyk, The influence of drug-polymer interactions on release of antirestenotic agent
from bioresorbable scaffolds, Materials Letters 223 (2018) 82-85.
[27] J. J. Hua, K. Gebarowska, P. Dobrzynski, J. Kasperczyk, J. Wei, S. M. Li, Influence of
chain microstructure on the hydrolytic degradation of copolymers from 1,3-trimethylene
carbonate and L-lactide, Journal of Polymer Science Part A-Polymer Chemistry 47 (15)
(2009) 3869-3879.
[28] J. Yang, F. Liu, L. Yang, S. M. Li, Hydrolytic and enzymatic degradation of
poly(trimethylene carbonate-co-D,L-lactide) random copolymers with shape memory
behavior, European Polymer Journal 46 (4) (2010) 783-791.
[29] Q. K. Guo, Z. Q. Lu, Y. Zhang, S. M. Li, J. Yang, In vivo study on the histocompatibility
and degradation behavior of biodegradable poly(trimethylene carbonate-co-D,L-lactide),
Acta Biochimica et Biophysica Sinica 43 (6) (2011) 433-440.
[30] L. Liao, J. T. Dong, L. Shi, Z. Y. Fan, S. M. Li, Z. Q. Lu, In vitro degradation behavior of
L-lactide/trimethylene carbonate/glycolide terpolymers and a composite with
poly(L-lactide-co-glycolide) fibers, Polymer Degradation and Stability 111 (2015)
203-210.
[31] H. L. Li, J. P. Chang, Y. Y. Qin, Y. Wu, M. L. Yuan, Y. J. Zhang,
Poly(lactide-co-trimethylene carbonate) and Polylactide/Polytrimethylene Carbonate
Blown Films, International Journal of Molecular Sciences 15 (2) (2014) 2608-2621.
[32] E. A. Bender, M. D. Adorne, L. M. Colome, D. S. P. Abdalla, S. S. Guterres, A. R.
Pohlmann, Hemocompatibility of poly(epsilon-caprolactone) lipid-core nanocapsules
stabilized with polysorbate 80-lecithin and uncoated or coated with chitosan,
International Journal of Pharmaceutics 426 (1-2) (2012) 271-279.
[33] S. Manju, K. Sreenivasan, Gold nanoparticles generated and stabilized by water soluble
curcumin-polymer conjugate: Blood compatibility evaluation and targeted drug delivery
onto cancer cells, Journal of Colloid and Interface Science 368 (2012) 144-151.
[34] F. Su, X. Shen, Y. F. Hu, V. Darcos, S. M. Li, Biocompatibility of thermo-responsive
PNIPAAm-PLLA-PNIPAAm triblock copolymer as potential drug carrier, Polymers for
Advanced Technologies 26 (12) (2015) 1567-1574.
[35] E. L. Zhang, H. Y. Chen, F. Shen, Biocorrosion properties and blood and cell
compatibility of pure iron as a biodegradable biomaterial, Journal of Materials
Science-Materials in Medicine 21 (7) (2010) 2151-2163.
[36] Y. Nie, Y. Duan, Z. Zhang, Evaluation of blood compatibility of surface modification
treated poly (D, L-lactic and glycolic acid) with mPEG block copolymer, Journal of
Biomedical Engineering 24 (2) (2007) 336-9.
[37] Z. Q. He, L. Z. Xiong, In-vitro Response of Fibroblasts to Poly-L-lactide Composite
Materials Containing Bovine Bone, Polymers & Polymer Composites 20 (6) (2012)
531-535.
[38] U. Krumsdorf, S. Ostermayer, K. Billinger, T. Trepels, E. Zadan, K. Horvath, H. Sievert,
Incidence and clinical course of thrombus formation on atrial septal defect and patient
foramen ovale closure devices in 1,000 consecutive patients, Journal of the American
College of Cardiology 43 (2) (2004) 302-9.
[39] L. Liao, J. T. Dong, G. X. Wang, Z. Y. Fan, S. M. Li, Z. Q. Lu, Microstructure-property
relationship of L-lactide/trimethylene carbonate/glycolide terpolymers as cardiovascular
stent material, European Polymer Journal 66 (2015) 429-436.
[40] X. Zhang, M. A. Geven, D. W. Grijpma, J. E. Gautrot, T. Peijs, "Polymer-polymer
composites for the design of strong and tough degradable biomaterials", Materials Today
Communications 8 (2016) 53-63.
[41] J. T. Dong, L. Liao, L. Shi, Z. S. Tan, Z. Y. Fan, S. M. Li, Z. Q. Lu, A bioresorbable
cardiovascular stent prepared from L-lactide, trimethylene carbonate and glycolide
terpolymers, Polymer Engineering and Science 54 (6) (2014) 1418-1426.
[42] J. T. Dong, L. Liao, Y. Ma, L. Shi, G. X. Wang, Z. Y. Fan, S. M. Li, Z. Q. Lu,
Enzyme-catalyzed degradation behavior of L-lactide/trimethylene carbonate/glycolide
terpolymers and their composites with poly(L-lactide-co-glycolide) fibers, Polymer
Degradation and Stability 103 (2014) 26-34.
[43] Z. Zhang, R. Kuijer, S. K. Bulstra, D. W. Grijpma, J. Feijen, The in vivo and in vitro
degradation behavior of poly(trimethylene carbonate), Biomaterials 27 (9) (2006)
1741-1748.
[44] Y. R. Han, Z. Y. Fan, Z. Q. Lu, Y. Zhang, S. M. Li, In vitro Degradation of
Poly[(L-lactide)-co-(trimethylene carbonate)] Copolymers and a Composite with
Poly[(L-lactide)-co-glycolide] Fibers as Cardiovascular Stent Material, Macromolecular
Materials and Engineering 297 (2) (2012) 128-135.
[45] J. Li, W. Zheng, Y. F. Zheng, X. Lou, Cell responses and hemocompatibility of
g-HA/PLA composites, Science China-Life Sciences 54 (4) (2011) 366-371.
[46] B. C. Dash, G. Rethore, M. Monaghan, K. Fitzgerald, W. Gallagher, A. Pandit, The
influence of size and charge of chitosan/polyglutamic acid hollow spheres on cellular
internalization, viability and blood compatibility, Biomaterials 31 (32) (2010) 8188-8197.
[47] Q. S. Zhu, L. B. Chen, P. Y. Zhu, J. F. Luan, C. Mao, X. H. Huang, J. Shen, Preparation of
PNIPAM-g-P (NIPAM-co-St) microspheres and their blood compatibility, Colloids and
Surfaces B-Biointerfaces 104 (2013) 61-65.
[48] Y. L. Luo, C. H. Zhang, F. Xu, Y. S. Chen, Novel THTPBA/PEG-derived highly branched
polyurethane scaffolds with improved mechanical property and biocompatibility,
Polymers for Advanced Technologies 23 (3) (2012) 551-557.
[49] M. Therin, P. Christel, S. Li, H. Garreau, M. Vert, In vivo degradation of massive poly
(α-hydroxy acids): validation of in vitro findings, Biomaterials 13 (1992) 594.
[50] J. Fernandez, A. Larranaga, A. Etxeberria, J. R. Sarasua, Effects of chain microstructures
and derived crystallization capability on hydrolytic degradation of
poly(L-lactide/epsilon-caprolactone) copolymers, Polymer Degradation and Stability 98
(2) (2013) 481-489.
Declaration of interest
On behalf of all authors, I declare that there are no Conflicts of Interest in this work.
Suming LI
July 2nd, 2019